In the field of X-ray detection it is well-known to employ so-called intensifying screens to increase the radiation available for detection purposes. Such screens contain an X-ray luminescent material (e.g., phosphor) which is selected to emit a relatively large number of light photons for each X-ray photon striking the material. This effectively amplifies the X-rays to be detected since both the X-rays themselves and light emitted by X-ray-induced emission from the luminescent material are available for detection on film or other detection mediums or devices, such as arrays of light-sensitive electronic sensors (e.g., photodiodes, photoconductors, charge-coupled devices, etc.). The primary incentive to use such intensifying screens in medical applications is to reduce the amount of X-ray radiation which is required to produce a given exposure, thereby reducing the radiation risk to which a patient or operator is exposed.
Detector panels envisioned for digital radiography ideally employ pixelized phosphor screens which are aligned with the electronic elements used to detect the photons generated when an X-ray interacts with the phosphor. In the use of digital radiographic screens for a majority of diagnostic X-rays, the pixel size that is required in order to obtain the necessary resolution is on the order of 85 microns. In the case of digital radiography for mammography, this pixel dimension decreases to the order of 40 microns. In any digital radiographic scheme which utilizes a pixelized phosphor to improve the resolution and isolation of adjacent pixels, the relative alignment of the two elements, the phosphor pixel and the detector pixel, becomes critical. Several schemes to eliminate this precise alignment of two components have centered around the direct formation of the phosphor pixels on the sensor pixels by the growth of a columnar phosphor over specific regions, or the cutting or ablating of phosphors deposited on sensing elements, into pixels.
An alternative approach to pixelizing directly on the sensor device is to separately pixelize both a phosphor sheet and a sensor array. The two pixelized arrays are then joined together while insuring that the pixels are precisely aligned for the two layers. For example, the method used in U.S. Pat. No. 5,153,438 utilizes a pixelized phosphor screen which is bonded to an array of detector elements. In this patent, the pixels of the phosphor are generated to be the same size and shape as the active area of the detector array, formed on a silicon wafer. A series of alignment marks are included in both the phosphor substrate and the detector array substrate to facilitate the precise alignment of the individual pixels of the phosphor with the pixels of the detectors that is required.
The need for precise alignment of the various elements of the radiographic panel is exemplified by the need for a radiologist interpreting the X-ray image to discern radiographic features on the image from the artifacts of the imaging process. It becomes especially critical when the image under examination is a standard full size 14 by 17 inch radiographic image. Current approaches to the construction of such panels require the arrangement of individual sensor arrays into a large scale panel. The sensing elements are composed of individual detector elements (e.g., photodiodes) arranged in arrays formed on single crystal silicon wafers. A typical dimension for these wafers is four inches in diameter. A regular square array is built on the wafer, and the edges are trimmed to permit the alignment of adjacent wafers edge-to-edge to form the ultimate full-size panel. Unless sufficient care is exercised, an inactive area exists at the line of juncture of two adjacent wafers, thereby resulting in a line image on the output. During the interpretation of the radiographic image any feature occurring along this line image is lost.
X-ray image sensors utilizing fiber optic faceplates are known in the art. See, for example, U.S. Pat. Nos. 4,910,405 and 5,079,423. In such apparatus, a solid-state image-sensing device is connected to the phosphor screen through the fiber optic faceplate. The phosphor screen converts X-rays into light and then the fiber optic faceplate transmits an optical image onto the input side of the solid-state image-sensing device which transduces the optical image into an electric signal. In such conventional X-ray image sensors, the fiber optic faceplate is used to prevent X-ray radiation damage to the solid-stage image-sensing device. As the foregoing patents indicate, the core glass of the optical fibers sometimes contains cerium oxide, lanthanum oxide, barium oxide, lead oxide, etc. Such materials help to increase the ability of the fiber optic faceplate to absorb incident X-rays, thereby further deceasing the damage to the solid-state image-sensing device.
U.S. Pat. No. 4,593,400 discloses a tapered fiber optic faceplate used in conjunction with a scintillator and a detector for X-rays in dental applications. The objective in this patent in using the tapered optical fibers is to decrease the size of an optical image projected onto a particular photodetector or sensor. PCT Publn. No. WO 91/15786 discloses a fiber optic beam-imaging apparatus with tapered plastic fibers. As with U.S. Pat. No. 4,593,400, the tapered optical fibers are used to produce a reduced-scale optical image.
U.S. Pat. No. 5,008,547 discloses a fiber optics device containing an obliquely cut end surface which transmits an optical image projected onto that surface to a second end face of the fiber optics device, thereby forming an image with a reduced surface area on the second end face.
U.S. Pat. No. 5,129,028 discloses an improved grid-free modular, large screen display. The improvement resides in the use of light-guides to eliminate apparent spacings between abutted display modules.
U.S. Pat. No. 5,144,141 discloses a radiation imaging device having a plurality of scintillator elements that are each optically coupled to a plurality of internal gain photoconductors. Each photodetector is electrically coupled to a respective detect-and-hold circuit which amplifies and stores the pulse generator by the photodetector. The stored pulses are sampled via a multiplexed switching arrangement to allow the stored signal from each detect-and-hold circuit to be processed to produce a digitized image signal which corresponds to the energy level of and location on the array of the detected incident radiation. The digitized imaging signal is then supplied to display memory and analysis equipment for the device.
Although the foregoing radiation detectors or imaging devices have been satisfactory for their intended use, improvements in the design and fabrication of radiation detectors are constantly sought and needed in the industry. For example, it would be desirable to have a radiation detector where individual pixel elements of the phosphor are aligned with individual pixel elements of the sensor without having to precisely position inactive areas of the phosphor with inactive areas of the detector, particularly when the size of the pixels becomes less than 85 microns. It would also be desirable to have a radiation detector with individual detector arrays (tiles or sub-modules) laid together into a large size radiographic panel without having inactive areas between adjacent tiles which would serve to limit the information accessible during the reading of an image from a radiographic panel.